OCT Probes
Optical coherence tomography apparatus are fairly well known and comprise a low coherent light source and an optical interferometer, commonly designed as either a Michelson optical fiber interferometer or a Mach-Zender optical fiber interferometer.
For instance, an optical coherence tomography apparatus known from the paper by X. Clivaz et al., “High resolution reflectometry in biological tissues”, OPTICS LETTERS, Vol. 17, No. 1, Jan. 1, 1992, includes a low coherent light source and a Michelson optical fiber interferometer comprising a beam-splitter optically coupled with optical fiber sampling and reference arms. The sampling arm incorporates an optical fiber piezoelectric phase modulator and has an optical probe at its end, whereas the reference arm is provided with a reference mirror installed at its end and connected with a mechanical in-depth scanner which performs step-by-step alteration of the optical length of this arm within a fairly wide range (at least several tens of operating wavelengths of the low coherent light source), which, in turn, provides information on microstructure of objects at different depths. Incorporating a piezoelectric phase modulator in the interferometer arm allows for lock-in detection of the information-carrying signal, thus providing a fairly high sensitivity of measurements.
The apparatus for optical coherence tomography reported in the paper by J. A. Izatt, J. G. Fujimoto et al., Micron-resolution biomedical imaging with optical coherence tomography, Optics & Photonics News, October 1993, Vol. 4, No. 10, p. 14-19 comprises a low coherent light source and an optical fiber interferometer designed as a Michelson interferometer. The interferometer includes a beam-splitter, a sampling arm with a measuring probe at its end, and a reference arm, whose end is provided with a reference mirror, movable at constant speed and connected with an in-depth scanner. This device allows for scanning the difference in the optical lengths of the sampling and reference arms. The information-carrying signal is received in this case using a Doppler frequency shift induced in the reference arm by a constant speed movement of the reference mirror.
Another optical coherence tomography apparatus comprising a low coherent light source and an optical fiber interferometer having a beam-splitter optically coupled to a sampling and reference arms is known from RU Pat. No. 2,100,787, dated 1997. At least one of the arms includes an optical fiber piezoelectric in-depth scanner, allowing changing of the optical length of said interferometer arm by at least several tens of operating wavelengths of the light source, thus providing information on microstructure of media at different depths. Since piezoelectric in-depth scanner is a low-inertia element, this device can be used to study media whose characteristic time for changing of optical characteristics or position relative to the optical probe is very short (the order of a second).
A major disadvantage inherent in all of the above-described apparatus as well as in other known apparatus of this type is that studies of samples in the direction approximately perpendicular to the direction of propagation of optical radiation are performed either by respective moving of the samples under study or by scanning a light beam by means of bulky lateral scanners incorporated into galvanometric probes. This does not allow these devices to be applied for medical diagnostics of human cavities and internal organs in vivo, as well as for industrial diagnostics of hard-to-access cavities.
Apparatus for optical coherence tomography known from U.S. Pat. No. 5,383,467, 1995 comprises a low coherent light source and an optical interferometer designed as a Michelson interferometer. This interferometer includes a beam-splitter, a sampling arm with an optical fiber sampling probe installed at its end, and a reference arm whose end is provided with a reference mirror connected with an in-depth scanner, which ensures movement of the reference mirror at a constant speed. The optical fiber sampling probe is a catheter, which comprises a single-mode optical fiber placed into a hollow metal tube having a lens system and an output window of the probe at its distal end. The optical tomography apparatus includes also a lateral scanner, which is placed outside the optical fiber probe and performs angular and/or linear scanning of the optical radiation beam in the output window of the optical fiber probe. However, although such geometry allows for introducing the probe into various internal cavities of human body and industrial objects, the presence of an external relative to the optical fiber probe lateral scanner and scanning the difference in the optical lengths of the sampling and reference arms by means of mechanical movement of the reference mirror significantly limit the possibility of using this device for performing diagnostics of surfaces of human cavities and internal organs in vivo, as well as for industrial diagnostics of hard-to-access cavities.
Apparatus for optical coherence tomography known from U.S. Pat. No. 5,582,171, 1996 comprises a low coherent light source and an optical fiber interferometer designed as a Mach-Zender interferometer having optical fiber sampling and reference arms and two beam-splitters. The reference arm includes a unit for changing the optical length of this arm. This unit is designed as a reference mirror with a spiral reflective surface arranged with a capability of rotating and is connected with a driving mechanism that sets the reference mirror in motion. The sampling arm is provided with an optical fiber probe having an elongated metal cylindrical body with a throughhole extending therethrough, and an optical fiber extending through the throughhole. A lateral scanner is placed at the distal end of the probe, which lateral scanner comprises a lens system, a rotatable mirror, and a micromotor for rotating the mirror, whereas an output window of the probe is located in the side wall of the cylindrical body. This device allows imaging of walls of thin vessels, but is unsuitable as a diagnostic means to image surfaces of cavities and internal organs inside a human body, as well as for industrial diagnostics of hard-to-access large-space cavities.
Another optical coherence tomography apparatus is known from U.S. Pat. No. 5,321,501, 1994 and comprises a low coherent light source optically coupled with an optical fiber Michelson interferometer, which includes a beam-splitter and optical fiber sampling and reference arms. The reference arm has a reference mirror mounted at its end and connected with an in-depth scanner. The latter performs movement of the reference mirror at a constant speed, thereby changing the optical length of this arm by at least several tens of operating wavelengths of the light source. The interferometer also comprises a photodetector whose output is connected with a data processing and displaying unit, and a source of control voltage connected with the in-depth scanner. The sampling arm incorporates an optical fiber probe having an elongated body with a throughhole extending therethrough, wherein a sheath with an optical fiber embedded in it extends through the throughhole. The sheath is attached to the stationary body through a pivot joint. The probe body contains also a lateral scanner comprising a bearing support, an actuator, and a lens system. The actuator includes a moving part and a stationary part, whereas the bearing support, the stationary part of the actuator and the lens system are mechanically connected with the probe body. The fiber-carrying sheath rests on the moving part of the actuator. The actuator may be a piezoelectric element, stepper motor, electromagnetic system or electrostatic system. The distal part of the probe body includes a lens system, the end face of the distal part of the optical fiber being optically coupled with the lens system, whereas the actuator is connected with a source of control current. The output of the data processing and displaying unit of the optical fiber interferometer is the output of the apparatus for optical coherence tomography. A disadvantage of this apparatus is that it is not fit for diagnostics of surfaces of hard-to-access internal human organs in vivo, such as, for example, stomach and larynx, and for industrial diagnostics of surfaces of hard-to-reach cavities of technical objects. That is due to the fact that the optical fiber probe in this apparatus must have relatively large dimensions since maximum movement of the optical fiber relative to the size of the actuator cannot be more than 20%, because of the moving part of the actuator being positioned at one side of the fiber-carrying sheath. Besides, the mechanical movement of the reference mirror at a constant speed used for scanning the difference in optical lengths of the reference and sampling arms restricts the range of objects, which can be studied in vivo by this apparatus, or by any other apparatus of this kind, to those objects whose optical characteristics and position relative to the optical probe do not change practically in the process of measurements.
Particular attention was also given to studies of biological tissues in vivo. For instance, a method for studying biological tissue in vivo is known from U.S. Pat. No. 5,321,501, 1994 and U.S. Pat. No. 5,459,570, 1995, in which a low coherent optical radiation beam at a given wavelength is directed towards a biological tissue under study, specifically ocular biological tissue, and to a reference mirror along the first and the second optical paths, respectively. The relative optical lengths of these optical beam paths are changed according to a predetermined rule; radiation backscattered from ocular biological tissue is combined with radiation reflected from a reference mirror. The signal of interference modulation of the intensity of the optical radiation, which is a result of this combining, is used to acquire an image of the ocular biological tissue. In a particular embodiment, a low coherent optical radiation beam directed to biological tissue under study is scanned across the surface of said biological tissue.
A method for studying biological tissue in vivo is known from U.S. Pat. No. 5,570,182, 1996. According to this method, an optical radiation beam in the visible or near IR range is directed to dental biological tissue. An image is acquired by visualizing the intensity of scattered radiation. The obtained image is then used for performing diagnostics of the biological tissue. In a particular embodiment, a low coherent optical radiation beam is used, which is directed to dental tissue, said beam being scanned across the surface of interest, and to a reference mirror along the first and second optical paths, respectively. Relative optical lengths of these optical paths are changed in compliance with a predetermined rule; radiation backscattered from the dental tissue is combined with radiation reflected by the reference mirror. A signal of interference modulation of intensity of the optical radiation, which is a result of said combining, is used to visualize the intensity of the optical radiation backscattered from said biological tissue. However, this method, as well as other known methods, is not intended for performing diagnostics of biological tissue covered with epithelium.
Acoustic Radiation Force Imaging
Ultrasound imaging is a non-invasive, diagnostic modality that is capable of providing information concerning tissue properties. In the field of medical imaging, ultrasound may be used in various modes to produce images of objects or structures within a patient. In a transmission mode, an ultrasound transmitter is placed on one side of an object and the sound is transmitted through the object to an ultrasound receiver. An image may be produced in which the brightness of each image pixel is a function of the amplitude of the ultrasound that reaches the receiver (attenuation mode), or the brightness of each pixel may be a function of the time required for the sound to reach the receiver (time-of-flight mode). Alternatively, if the receiver is positioned on the same side of the object as the transmitter, an image may be produced in which the pixel brightness is a function of the amplitude of reflected ultrasound (reflection or backscatter or echo mode). In a Doppler mode of operation, the tissue (or object) is imaged by measuring the phase shift of the ultrasound reflected from the tissue (or object) back to the receiver.
Ultrasonic transducers for medical applications are constructed from one or more piezoelectric elements activated by electrodes. Such piezoelectric elements may be constructed, for example, from lead zirconate titanate (PZT), polyvinylidene diflouride (PVDF), PZT ceramic/polymer composite, and the like. The electrodes are connected to a voltage source, a voltage waveform is applied, and the piezoelectric elements change in size at a frequency corresponding to that of the applied voltage. When a voltage waveform is applied, the piezoelectric elements emit an ultrasonic wave into the media to which it is coupled at the frequencies contained in the excitation waveform. Conversely, when an ultrasonic wave strikes the piezoelectric element, the element produces a corresponding voltage across its electrodes. Numerous ultrasonic transducer constructions are known in the art.
When used for imaging, ultrasonic transducers are provided with several piezoelectric elements arranged in an array and driven by different voltages. By controlling the phase and amplitude of the applied voltages, ultrasonic waves combine to produce a net ultrasonic wave that travels along a desired beam direction and is focused at a selected point along the beam. By controlling the phase and the amplitude of the applied voltages, the focal point of the beam can be moved in a plane to scan the subject. Many such ultrasonic imaging systems are well known in the art.
An acoustic radiation force is exerted by an acoustic wave on an object in its path. The use of acoustic radiation forces produced by an ultrasound transducer has been proposed in connection with tissue hardness measurements. See Sugimoto et al., “Tissue Hardness Measure Using the Radiation Force of Focused Ultrasound”, IEEE Ultrasonics Symposium, pp. 1377-80, 1990. This publication describes an experiment in which a pulse of focused ultrasonic radiation is applied to deform the object at the focal point of the transducer. The deformation is measured using a separate pulse-echo ultrasonic system. Measurements of tissue hardness are made based on the amount or rate of object deformation as the acoustic force is continuously applied, or by the rate of relaxation of the deformation after the force is removed.
These and other documents may provide additional context where necessary for fuller understanding of the claimed invention and are incorporated by reference herein in their entirety for references purposes and for assisting in the determination of the level of ordinary skill in the art, e.g. U.S. Pat. No. 7,022,077.